In general, an aspect of the present invention relates to the field of magnetic resonance (MR) in which appropriate magnet systems are employed, which are designed to generate uniform magnetic fields for MR measurements. However, the applicability of the invention is not limited to this field.
In both nuclear magnetic resonance spectroscopy (NMR) and in imaging applications (MRI), within a specific sample volume, an exceptionally uniform and temporally constant magnetic field is required, which can be generated by resistive or superconducting coils, or by an appropriate permanent magnet arrangement. Magnetic resonance imaging (MRI) methods are extensively employed for the acquisition of image information on structures. Using these methods, it is possible to obtain image information from the interior of a structure without damaging said structure. In clinical applications, for example, internal images of the organs of human beings and animals can be generated using MRI methods. In currently employed MRI methods, precision movements of nuclear spin are characterized by the temporally-varied superimposition of additional location-related magnetic fields for all three spatial directions in a spatial coding system, generally described as spatial encoding. Within the investigated object, these additional magnetic fields customarily show essentially constant gradients for the z-components in the spatial directions x, y and z, and are generated by a coil arrangement which is designated as a gradient system, and are actuated by a “gradient channel” for each respective spatial direction.
The generation of a gradient is achieved in that a pulse-shaped current is delivered by a gradient coil, which varies the static magnetic field on the investigated object. As described, for example, in US 2006/0152222 A1, an increased voltage, of up to several hundred volts, is applied to the gradient coil, as a result of which the current rises more rapidly such that, once the desired current strength is achieved, a switchover to a lower voltage is effected, in order to reduce the power loss in the gradient amplifier. In the first case, however, a high voltage is required, albeit for only a short time, which must still be available in reserve in the gradient amplifier circuit.
In electrotechnical circuit theory and network analysis, a current source describes an active two-terminal circuit which delivers an electric current at its terminal points. As an important property, this current is only dependent to a limited extent or, in the model of an ideal current source in the context of circuit analysis, is not at all dependent upon the electrical voltage at the terminal points thereof.
In electrotechnical circuit theory, a voltage source describes an active two-terminal circuit which delivers an electrical voltage between its terminal points. As an important property, this voltage is only dependent to a limited extent or, in the model of an ideal voltage source in the context of network analysis, is not at all dependent upon the electric current which is tapped from the source.
The above-mentioned US 2006/0152222 A1 discloses a gradient circuit with an amplifier circuit of conventional operation, wherein the gradient amplifier comprises a plurality of switching current controllers, which are electrically connected in series. A bipolar circuit receives power from the associated series circuit power controller and delivers power, with the preferred polarity, to the gradient coil.
DE 33 36 286 A1 constitutes an alternative to the arrangement described in US 2006/0152222 A1. However, the current source employed therein, which has a high output resistance and delivers a constant current, according to the definition of a current source, is not appropriate as an infeed source for the generation of an optimum square-wave rising current ramp in the gradient coil. In the circuit described, current from the current source initially flows through an auxiliary coil, as a result of which a magnetic field is constituted therein. When the commencement of the gradient current is required, the latter is switched-on and the current in the auxiliary coil is switched-off. A high voltage is present on the auxiliary coil for a short time. As a result of the employment of the current source as a current supply, however, the energy stored in the auxiliary coil cannot flow into the gradient coil, as the output resistance of the current source is high. Charging of the gradient coil by an auxiliary inductance is not described in US 2006/0152222 A1.
A further alternative is constituted by the circuit described in U.S. Pat. No. 4,961,054. In this case, the required gradient current is supplied by an amplifier. As this current is evidently maintained at a constant value, an auxiliary coil with a 5 to 20 times higher inductance than that of the gradient coil is required. With an inductance of this magnitude, it is a very long time before the requisite current flux can be achieved, or the circuit which supplies the amplifier requires a very high maximum output voltage. Thus, in the measuring system, either a very high supply voltage for the amplifier is required, or the waiting time between two gradient pulses must be relatively long. Even in the presence of the above-mentioned conditions, it is not possible to generate an optimum pulse-shaped rising current in the gradient coil. This would require an auxiliary coil of infinite inductance, resulting in an infinite waiting time, or an amplifier with an infinite maximum output voltage. Neither is possible in practice.
U.S. Pat. No. 5,270,657 further describes a gradient system in which DC current supplies are connected in tandem with conventional linear gradient amplifiers, in order to increase the effective gradient energy supplied to the gradient coils. Here again, charging of the gradient coils by an auxiliary inductance is not described.
Finally, U.S. Pat. No. 2,941,125 discloses a circuit for the generation of currents with an optimum square-wave profile in an inductance, for the switching of a microwave device.